The present invention generally relates to use of ultrasound for imaging and therapeutic purposes, and more specifically, to simplified ultrasound transducers that are both highly efficient for administering therapy and produce a wide bandwidth ultrasound signal for diagnostic imaging.
Ultrasound is a term that refers to acoustic waves having a frequency above the upper limit of the human audible range (i.e., above 20 kHz). Because of their relatively short wavelength, ultrasound waves are able to penetrate into the human body. Based on this property, ultrasound in the frequency range of 2-20 MHz has been widely used to image internal human organs for diagnostic purposes.
To avoid thermal damage to tissue, the power level in diagnostic ultrasound imaging is kept very low. The typical ultrasound intensity (power per unit area) used in imaging is less than 0.1 watt per square centimeter. High intensity focused ultrasound, which can have an intensity above 1000 watts per square centimeter, can raise the tissue temperature at the region of the spatial focus to above 60 degrees Celsius in a few seconds and can cause tissue necrosis almost instantaneously.
High intensity ultrasound has been proposed to treat and destroy tissues in the liver (G. ter Haar, xe2x80x9cUltrasound Focal Beam Surgery,xe2x80x9d Ultrasound in Medicine and Biology, Vol. 21, No. 9, pp.1089-1100, 1995); in the prostate (N. T. Sanghvi and R. H. Hawes, xe2x80x9cHigh-intensity Focused Ultrasound,xe2x80x9d Experimental and Investigational Endoscopy, Vol. 4, No. 2, pp.383-395, 1994); and in other organs.
Ultrasound transducers generate ultrasound waves for imaging and therapy. A typical ultrasound transducer comprises piezoelectric materials such as PZT ceramics, electrodes, matching layers, and backing materials. When an electrical field is applied to two electrodes on the opposite sides of a piezoelectric ceramic plate, the thickness of the plate expands or contracts, depending on the polarity of the field. If the electrical field polarity alternates at a high frequency above 20 kHz, the mechanical vibration caused by the rapid expansion/contraction of the plate generates ultrasound waves.
During ultrasound therapy, high electrical power is applied to the ultrasound transducer to generate a correspondingly high acoustical output power. Transducer power conversion efficiency is the ratio of the output acoustic power to the input electrical power. A high transducer power conversion efficiency is always desirable to minimize the transducer internal heating due to electrical power losses.
During ultrasound imaging, low-power electrical pulses drive the transducer, causing it to transmit the low power ultrasound pulses into the patient body. Ultrasound echoes, reflected from organ boundaries and other tissue and physiological structures within the body, are typically received by the same ultrasound transducer and converted to electrical output signals, which are processed to produce ultrasound images of the internal organ on a display. A transducer having a broad frequency bandwidth is desirable to obtain good image resolution. Often, however, the desire for high efficiency during ultrasound therapy and the desire for broad bandwidth during ultrasound imaging are difficult to satisfy simultaneously in the same transducer design.
To treat or to image a large volume of diseased tissue, the ultrasound beam is caused to scan through the tissue, either mechanically or electronically. In a mechanical scanning device, such as disclosed in U.S. Pat. No. 4,938,216, one or more electrical motors position the ultrasound transducer in different positions. One of the more common types of electronic scanning device employs an ultrasound linear phased-array transducer, such as that disclosed by E. B. Hutchinson and K. Hynynen, in an article entitled xe2x80x9cIntracavitary Ultrasound Phased Arrays for Noninvasive Prostate Surgeryxe2x80x9d (IEEE Transactions on Ultrasonics, Ferroelectrics, and Frequency Control, vol. 43, No. 6, pp.1032-1042, (1996)), and in U.S. Pat. No. 4,938,217. An electronic scanning device has a plurality of small piezoelectric elements disposed in an array. These elements are independently driven. By properly controlling the phase of the driving signals applied to energize these elements, the array will be caused to form ultrasound beams directed at different depths and angles. The electronic scanning transducer has many advantages over the mechanically scanned transducer. The main advantage is that there are no moving components in the electronic device, so that it has much higher durability and reliability. The disadvantage of the electronic device is its complexity and associated relatively high cost. To achieve a compromise between the advantages and the disadvantages, some prior art references, such as U.S. Pat. No. 4,757,820, disclose transducer designs that include both the mechanical and the electronic approaches.
However, system and transducer complexity is still one of the major disadvantages of electronic therapeutic arrays. A therapeutic transducer requires a large surface area to generate a high acoustic power output and a large aperture for deep treatment. Preferably, the f-number (focal depth over aperture size) is kept constant within the range from 0.8 to 2.5. On the other hand, to steer the ultrasound beam over a wide range and to focus the beam using a small f-number, the ultrasound phased array must have very fine, narrow array elements, because a narrow element can transmit an ultrasound beam throughout a wide range of directions.
To provide a transducer having both a large aperture and fine elements to enable it to provide imaging and therapy functions, a conventional therapeutic phased array design includes a very large number of elements. For example, to treat lesions at a maximum depth of 5 cm, a therapeutic linear array having an f-number of 1.0 should have an aperture width of about 5 cm. For use at this depth, the transducer will typically operate at a frequency of about 3 MHz. The wavelength of ultrasound in water or biological soft tissue at this frequency is about 0.5 mm. For a phased array of this configuration to have a sharp focus (i.e., a relatively small f-number), the array typically will have an element pitch size about 0.5 to 0.7 times the wavelength of the ultrasound beam it produces. For a pitch size of about 0.6 times the wavelength, an exemplary therapeutic array might have an element pitch size of about 0.3 mm and a total of about 167 elements.
Each element has a dedicated electronic driving circuit in a control system for the array. To drive a phased array like that discussed above, the control system would need to include 167 sets of driving circuits, i.e., one for each element. The array and the control system are connected through a thick cable that includes at least 167 smaller coaxial cables inside it. Each smaller coaxial cable should have conductors of a sufficiently large cross-sectional area to carry a relatively large current to the therapeutic array element. The thick cable required to meet this need makes the device difficult to handle.
Considering all these constraints, it will be evident that the complexity of such a therapeutic phased array, including the cable and the control system coupled to it, can easily become impractical to engineer, and its cost will most certainly exceed the budget of most medical facilities. It is for these reasons that the therapeutic phased array has not been widely accepted.
It would be desirable to use an ultrasound array transducer for both imaging and therapy. The smaller size of a probe having a transducer that is usable for both functions is an advantage. For example, in many endoscopic, therapeutic-ultrasound applications, there are limitations on the size of the treatment devices that can be employed. Thus, a dual-purpose ultrasound array transducer may save space in the probe. Also, in ultrasound image-guided therapeutic applications, there are two spatial planes, one for imaging and the other for treatment. These two planes should overlap so that the treatment area can be observed in the imaging plane. Oftentimes, however, it is difficult to register the two planes from two spaced-apart transducers. Sometimes, there are blind spots in the treatment zone, which are not observable in the imaging plane. However, if one transducer is used both for imaging and treatment, the problem of non-overlapping zones does not arise.
The prior art has not dealt extensively with the problem of designing a dual-purpose phased array transducer. Besides the conflict between the disparate design parameters that must be satisfied to achieve efficiency and adequate bandwidth in such a transducer, as noted above, there are other unresolved issues in making a therapeutic phased array transducer, such as heat dissipation, and element cross-talk. In U.S. Pat. No. 6,050,943 and in an article published by P. G. Barthe and M. H. Slayton, entitled xe2x80x9cEfficient Wideband Linear Arrays for Imaging and Therapyxe2x80x9d (IEEE Symposium in Ultrasonics, Ferroelectrics and Frequency Control, November 1999), the authors address some of these problems.
Thus, there is a clear need for an ultrasound device that employs simple and highly efficient ultrasound transducer arrays usable for both imaging and therapy. This kind of ultrasound device can be used to generate real-time ultrasound images of a patient""s internal condition, provide ultrasound therapy to a treatment site, and monitor the treatment results. Such an ultrasound transducer should have variable geometry for treating different pathologies. In addition, the transducer array should be capable of generating high-intensity ultrasound to ablate or necrose tumors and other diseased tissues.
The present invention provides an ultrasound transducer apparatus comprising a generally concave array of ultrasound transducer elements. The apparatus enables a reduced number of transducer elements and a larger pitch size compared to that used for the elements in a traditional linear array of transducer elements. Reducing the number of elements also reduces the required number of connection cables and control channels. While providing the same performance, the concave array system is much simpler and less costly than a conventional linear phased array system. The concave geometry also requires smaller phase differences between transducer elements, thus reducing cross-talk and heating in kerf fills between elements. The geometry also reduces the affect of grating lobe problems during the beam-forming process.
To provide both imaging and therapy functions, one embodiment of the present invention includes circuitry to rapidly switch between low and high Q-factors. Alternatively, the invention may include one transducer array for imaging and another transducer array for therapy, enabling one of the arrays to selectively act on a target site. For example, the imaging transducer array and therapeutic transducer array may be attached to opposite sides of a rotatable carriage and alternately directed to the target site as the carriage rotates.
To control a location of a focus point of the transducer array, one form of the invention includes a beam steering mechanism, or controller, to adjust the phases or the delays of signals that drive the transducer elements. To increase the transducer bandwidth for better image resolution, an electrical damping circuit can be included to provide the equivalent of a mechanical backing. One or more material acoustic matching layers and/or air backing can optionally be included to improve the transducer efficiency and bandwidth. In addition, the present invention may optionally include one or more metal matching layers to improve heat dissipation by the transducer.
A flexible transducer array is preferably provided to control the location of the focus point. Flexible outer layers and kerf fills between transducer elements enable the array to bend in different curvatures. As with a fixed curvature array, the flexible array reduces the number of required transducer elements. However, the flexible array embodiment also enables a practitioner to adjust the imaging field of view (FOV) and simplifies control of the treatment focusing, by changing the geometric shape of the array.
To facilitate these capabilities, the invention may include a geometry control mechanism. Preferably, the control mechanism and flexible transducer array comprise a laparoscopic applicator in which a linear actuator translates one end of the flexible transducer array relative to an opposite fixed end, causing the transducer array to flex into a desired curved shape. The actuator alternatively comprises either a manual adjustable shaft or a motor-driven threaded shaft, shuttle block, push rod, or the like. Another embodiment includes position stops or a position template to guide the curvature of the array, so that the array matches the profile of the position stops or template. The position stops or template may be preset, or adjustable. The geometry control mechanism may also be independently applied to one transducer array that is dedicated to one of the functions of imaging or therapy, while another transducer array is dedicated to the other function. For example, in a laparoscopic applicator, the control mechanism may be applied to a therapy transducer array connected to a rotational carriage, while an imaging transducer array is attached to the opposite side of the rotational carriage and is not provided with any control mechanism.
Another embodiment of the invention includes a plurality of transducer arrays, each directed toward a common focus point. Using multiple transducer arrays enables each array to contain fewer transducer elements and provides a relatively wide imaging and treatment field. Each transducer array may also be allowed to pivot about a pivot point, such that controlled pivoting of the multiple transducer arrays controls the location of the common focus point. This enables controlled movement of the common focus point in at least two directions.
Another aspect of the invention includes a transducer manufacturing method to produce an ultrasound transducer apparatus with a generally concave geometry. The method comprises the step of providing kerf fills having a non-uniform stiffness to control the curvature of the transducer array. For example, providing kerf fills with a symmetrically non-uniform stiffness improves the likelihood of obtaining a symmetric semi-circle shape of the array, rather than a parabolic shape, when moving one end of a transducer array, compared to an array that has uniformly stiff kerf fills. Alternatively, or in addition, the method may include the step of providing support layers having a non-uniform stiffness. Another step of the method preferably includes cutting grooves into a metal support layer between the transducer elements on the side of the support layer that supports the transducer elements to avoid bonding between the transducer element and the metal support layer. Further steps optionally include cutting grooves into an opposite side of the support layer, and casting an outer matching layer over the support layer and into the grooves to improve the bonding strength between the support layer and an outer matching layer. If the outer matching layers, or support layers, are not deformable, alternate steps to provide flexibility include cutting the outer matching layer into thin strips after bonding the outer matching layer to the support layer, and then filling the kerfs with deformable material.